1. Field of the Invention
This invention pertains generally to radiation dose computation and more particularly to radiation dose computation for CT imaging.
2. Description of Related Art
Computed Tomography (CT) is a medical imaging procedure which is performed millions of times each year in the United States. CT scans (formerly called CAT scans) allow physicians to look at images of the anatomy of the patient in individual “slices” or “sections”. A modern CT scanner can produce upwards of 1,000 images of a patient's body part to be scanned, typically the head, thorax, abdomen, or pelvis.
Because of the enormous utility of computed tomography, CT has become the single largest contributor to radiation dose in the U.S. population. The radiation dose associated with computed tomography has always been of concern, however the increased use of CT across a broad spectrum of diagnostic situations coupled with the high output capabilities of modern CT scanners heightens these concerns.
Modern CT scanners have modes of operation which current CT dosimetry techniques do not and cannot address. In particular, the existence of both angular and z-axis mA modulation (discussed later) results in the output of a scanner being automatically changed during the actual scan, based on the patient's contour. Cylindrical Lucite phantoms are the current basis for the vast majority of CT dosimetry in the United States and beyond. These phantoms do not change in thickness as a function of either angle or z dimension, being homogenous polymethacrylate (PMMA) cylinders.
In addition to the inability to handle current scanner technology for accurate dosimetry, current CT dosimetry standards typically underestimate the radiation dose to the patient for a number of reasons. These reasons include the fact that most patients are not 32 cm in diameter, and patients' tissues have a density closer to 1.0 rather than the 1.19 gm/cm3 density of PMMA. The thicker, denser dosimetry phantoms lead to an underestimate of the central radiation doses associated with most body CT procedures.
The radiation dose in CT comes from two sources: the primary radiation dose (the deposition of energy in tissues from x-ray photons which started their trajectory in the x-ray tube focal track) and from scattered radiation (x-rays which have been scattered within the patient, and are redistributed, contributing dose appreciably to tissues away from the original x-ray interaction point in the patient). Current dosimetry techniques for CT try to accommodate the scatter aspects of dose by using a long (100 mm) so-called pencil chamber which is exposed while placed in a hole in a PMMA cylinder. Unfortunately, at the high energies used in x-ray CT, the scattered radiation propagates to distances in the z dimension of the phantom farther than what the 100 mm pencil chamber can measure. This geometry therefore leads to an underestimate of the radiation dose due to CT.
Thus, for a number of reasons, it is clear that the methodology for computing radiation dose to patients in CT procedures needs significant improvement to obtain the accuracy necessary to make informed decisions about risk assessment, and whether a patient should have additional or continuing CT for other x-ray procedures.
CT became a commercial product in the early 1970s, and since that time, the utility of the CT procedure has increased in importance with its growing capabilities and shorter scan times of CT scanners. A CT scanner acquires the raw data necessary for producing the CT images. As shown in FIG. 1, most modern CT scanners 10 have a x-ray tube 12 rotates in unison with the detector arrays 14 around the patient's body 16, emitting x-ray photons which interrogate the patient, and some of them emerge from the opposite side of the patient and strike the x-ray detector. The x-ray tube emits a number of photons per unit area (on the detector) No, but behind the patient, this value is reduced to N, due to x-ray attenuation from the patient's tissue. For a given path in the patient of distance X, the linear attenuation coefficient is computed by the scanner hardware using the following equation.N=N0e−μx  Equation 1
Where μ is the average linear attenuation coefficient of the patient along that path X.
Once the CT acquisition is complete, the acquired data is pre-processed and then used to produce the CT images in a procedure known as filtered back projection. After the CT images are reconstructed, the images are re-scaled depictions (two dimensional maps 18 as shown in FIG. 2) of the linear attenuation coefficient (LAC) distribution in the patient. The rescaled values of the LAC are called Hounsfield Units (HU).
Once the CT images are produced from the filtered back projection, cone beam reconstruction, or iterative reconstruction procedure they comprise a volume data set which essentially digitizes the patient into millions of individual volume elements (voxels).
Referring now to FIG. 3, CT scans are produced along a length of the patient's body, and a series of contiguous, or even overlapping, CT images 18 are produced to generate the volume data set 20. The series of images represents a map of the LAC of the patient, in three dimensions. The images are defined in the (x,y) plane, and the long axis of the scan is typically referred to as the z-axis. A series of axial images, which are acquired contiguously, form this volume data set. Thus, an (x,y,z) data set of linear attenuation coefficient values are the typical result of clinical CT scanning.
Each image in CT is a quantitative representation of the x-ray attenuation properties of an individual voxel in the patient. However, the CT scanner computer normalizes the reconstructed linear attenuation coefficient into a gray scale value of Hounsfield units (HU).
                    HU        =                  1000          ⁢                                    (                              μ                -                                  μ                  w                                            )                                      μ              w                                                          Equation        ⁢                                  ⁢        2            where HU is the Hounsfield Unit (gray scale of the CT image) of a given pixel in the image, μ is the linear attenuation coefficient (LAC) of the corresponding voxel in the patient, and μw is the linear attenuation coefficient of water for that scanner and for that x-ray beam.
A unique but key feature relative to the invention described here in CT is that the gray scale values (Hounsfield units) are quantitatively accurate descriptors of the x-ray attenuation properties of each voxel within the patient. The current method for computing x-ray dose to the patient relies upon this quantitative nature of the CT image data.
Referring now to FIG. 4, to compute doses in CT, conventional techniques use a pencil chamber 30 (a long thin cylindrical ionization chamber) placed in various holes in a PMMA phantom 32. The phantom 32 is scanned in a CT scanner, and the dose received by the chamber is recorded.
The Computed Tomography Dose Index (CTDI) has been defined in the Code of Federal Regulations (21-CFR-1020.33), where:
                    CTDI        =                              1            nT                    ⁢                                    ∫                                                -                  7                                ⁢                T                                                              +                  7                                ⁢                T                                      ⁢                                          D                ⁡                                  (                  z                  )                                            ⁢                              ⅆ                z                                                                        Equation        ⁢                                  ⁢        3            and where:
z is the position along the z-axis of the scan or patient,
D(z) is the dose at position z,
T is the nominal tomographic section thickness, and
n is the number of images produced in a single scan.
This assumes that the scan increment (between images) is nT.
Due to the standard 100 mm length of the pencil chamber used almost worldwide for CT dosimetry, the CTDI100mm has been defined as:
                              CTDI                      100            ⁢            mm                          =                              1            nT                    ⁢                                    ∫                                                -                  50                                ⁢                mm                                                              +                  50                                ⁢                mm                                      ⁢                                          D                ⁡                                  (                  z                  )                                            ⁢                              ⅆ                z                                                                        Equation        ⁢                                  ⁢        4            
The CTDI100mm is the basis for almost all CT dosimetry performed worldwide, with the exception of various research studies that use more sophisticated techniques such as MOSFET (metal oxide semiconductor field-effect transistor) dosimeters, thermoluminescent dosimeters (TLDs), or other radiation monitoring devices. These techniques are not practical for routine patient dosimetry, due to the time that they require, and the fact that measurement devices need to be placed internal to the subject being scanned—not feasible for live human imaging.
Although the CTDI was never intended by its originators as a direct measure of patient dose, over the years scientists and CT practitioners have sought to make the CTDI a dosimetric quantity. The CTDI100mm values can be measured on a specific scanner at both the center hole and the peripheral hole on the standard PMMA phantom (shown in FIG. 4). The CTDI100mm measurement at the center is dubbed CTDIcenter, and the peripheral measurement is called CTDIperiphery. The weighed CTDI, CTDIw, has been defined as:
                              CTDI          w                =                                            1              3                        ⁢                          CTDI              center                                +                                    2              3                        ⁢                          CTDI              peripheral                                                          Equation        ⁢                                  ⁢        5            
The CTDIw is thought to be a more accurate description of patient dose than CTDI100mm per se.
For multiple detector array helical CT scanners, which are the norm in modern CT facilities, the patient table is moved at constant velocity during the rotation of the CT gantry (the gantry consists of the x-ray tube, detector arrays, and other components on a rotate/rotate system). For a multiple detector array with n detector arrays, and a section thickness of T (per detector array), the table will translate a distance nT if the pitch is unity. For a table translation of S mm per complete rotation of the gantry (around 360 degrees or 2π radians), the pitch is defined as:
                    pitch        =                  s          nT                                    Equation        ⁢                                  ⁢        6            
The units of s and T, in equation 6, are in mm (or cm). When pitch<1, the dose to the patient increases because the x-ray beam over-samples the patient, and when pitch>1, the dose to the patient decreases because the x-ray beam under-samples the patient. When the pitch=1, the dose in helical CT scanning is almost the same as it is in contiguous axial scanning (which is the assumption in CTDIFDA, mentioned above in equation 3). To adjust dose to accommodate different pitch values used on the CT scanner, the “CTDI volume” has been defined as:
                              CTDI          vol                =                              CTDI            w                    pitch                                    Equation        ⁢                                  ⁢        7            
CTDIvol is meant to estimate the dose from a specific scan geometry (one slice), and this metric therefore does not take into consideration the fact that in clinical CT scanning, a length of the patient is typically scanned. The length of the CT scan in the body is often 30 to 50 cm, depending on the size of the patient and the body region to be scanned. To account for the dose in these longer scans, the dose length product (DLP) has been defined as:DLP=CTDIvol×scan_length  Equation 8
The unit of DLP is not even a dose unit, but rather has the units of (mGy cm). The DLP is often displayed on the CT console during the scan, as a very crude and scientifically obscure metric for the radiation dose that the patient being scanned receives.
There are number of developments in CT technology which suggest that a homogeneous PMMA cylinder is no longer adequate (it never was accurate) for patient dosimetry. Specifically, CT scanners are now capable of changing the radiation output of the x-ray tube as the tube rotates (θ mA modulation) and as the patient table is translated (z-axis mA modulation).
Referring now to FIG. 5, as the tube rotates around the typical patient, depending on the location in the body, the profile of the patient is usually elliptical and not circular. To obtain the best image quality at the lowest radiation dose levels, modern CT scanners turn down the output of the x-ray tube at location A, where the projection of the patient is thinner and less radiation is needed, but turn up the radiation output of the tube at location B, where more x-rays are needed to penetrate the thicker patient projection at that angle.
Referring now to FIG. 6, Z-axis modulation works by changing the output of the x-ray tube during the CT scan as the effective thickness of the patient 16 changes. For example, less radiation is needed to penetrate the lung fields 34 (due to the low density of the lungs), while higher radiation levels are needed to penetrate the thicker abdomen 36.
Interestingly, both the θ and z-axis mA modulation schemes are determined by the shape (x-ray transmission properties) of the patient. The PMMA cylinder used for conventional CT dosimetry is constant in both θ and z, and does not represent an actual patient's shape. Thus, without knowing the patient's shape, accurate dosimetry would be virtually impossible.